Radiographic imaging such as x-ray imaging has been used for years in medical applications and for non-destructive testing.
Normally, an x-ray imaging system includes an x-ray source and an x-ray detector system. The x-ray source emits x-rays, which pass through a subject or object to be imaged and are then registered by the x-ray detector system. Since some materials absorb a larger fraction of the x-rays than others, an image is formed of the subject or object.
It may be useful to begin with a brief overview of an illustrative overall x-ray imaging system, with reference to FIG. 1. In this non-limiting example, the x-ray imaging system 100 basically comprises an x-ray source 10, an x-ray detector system 20 and an associated image processing device 30. In general, the x-ray detector system 20 is configured for registering radiation from the x-ray source 10 that may have been focused by optional x-ray optics and passed an object or subject or part thereof. The x-ray detector system 20 is connectable to the image processing device 30 via suitable analog processing and read-out electronics (which may be integrated in the x-ray detector system 20) to enable image processing and/or image reconstruction by the image processing device 30.
A challenge for x-ray imaging detectors is to extract maximum information from the detected x-rays to provide input to an image of an object or subject where the object or subject is depicted in terms of density, composition and structure. It is still common to use film-screen as detector but mostly the detectors today provide a digital image.
Modern x-ray detectors normally need to convert the incident x-rays into electrons, this typically takes place through photo absorption or through Compton interaction and the resulting electrons are usually creating secondary visible light until its energy is lost and this light is in turn detected by a photo-sensitive material. There are also detectors, which are based on semiconductors and in this case the electrons created by the x-ray are creating electric charge in terms of electron-hole pairs which are collected through an applied electric field.
There are detectors operating in an integrating mode in the sense that they provide an integrated signal from a multitude of x-rays and the signal is only later digitized to retrieve a best guess of the number of incident x-rays in a pixel.
Photon counting detectors have also emerged as a feasible alternative in some applications; currently those detectors are commercially available mainly in mammography. The photon counting detectors have an advantage since in principal the energy for each x-ray can be measured which yields additional information about the composition of the object. This information can be used to increase the image quality and/or to decrease the radiation dose.
Compared to the energy-integrating systems, photon-counting CT has the following advantages. Firstly, electronic noise that is integrated into signal by the energy-integrating detectors can be rejected by setting the lowest energy threshold above the noise floor in the photon-counting detectors. Secondly, energy information can be extracted by the detector, which allows improving contrast-to-noise ratio by optimal energy weighting and which also allows so-called material basis decomposition, by which different components in the examined patient can be identified and quantified, to be implemented effectively. Thirdly, more than two basis materials can be used which benefits decomposition techniques, such as K-edge imaging whereby distribution of contrast agents, e.g. iodine or gadolinium, are quantitatively determined. Fourth, there is no detector afterglow, meaning that high angular resolution can be obtained. Last but not least, higher spatial resolution can be achieved by using smaller pixel size.
The most promising materials for photon-counting x-ray detectors are cadmium telluride (CdTe), cadmium zinc telluride (CZT) and silicon (Si). CdTe and CZT are employed in several photon-counting spectral CT projects for the high absorption efficiency of high-energy x-rays used in clinical CT. However, these projects are slowly progressing due to several drawbacks of CdTe/CZT. CdTe/CZT have low charge carrier mobility, which causes severe pulse pileup at flux rates ten times lower than those encountered in clinical practice. One way to alleviate this problem is to decrease the pixel size, whereas it leads to increased spectrum distortion as a result of charge sharing and K-escape. Also, CdTe/CZT suffer from charge trapping, which would lead to polarization that causes a rapid drop of the output count rate when the photon flux reaches above a certain level.
In contrast, silicon has higher charge carrier mobility and is free from the problem of polarization. The mature manufacturing process and comparably low cost are also its advantages. But silicon has limitations that CdTe/CZT does not have. Silicon sensors must accordingly be quite thick to compensate for its low stopping power. Typically, a silicon sensor needs a thickness of several centimeters to absorb most of the incident photons, whereas CdTe/CZT needs only several millimeters. On the other hand, the long attenuation path of silicon also makes it possible to divide the detector into different depth segments, as will be explained below. This in turn makes it possible for a silicon-based photon-counting detector to properly handle the high fluxes in CT.
When using simple semiconductor materials, as silicon or germanium, Compton scattering causes many x-ray photons to convert from a high energy to a low energy before conversion to electron-hole pairs in the detector. This results in a large fraction of the x-ray photons, originally at a higher energy, producing much less electron-hole pairs than expected, which in turn results in a substantial part of the photon flux appearing at the low end of the energy distribution. In order to detect as many of the x-ray photons as possible, it is therefore necessary to detect as low energies as possible.
Indeed, photon counting x-ray imaging has gained considerable attention the last decade and in some cases matured into clinical applications. In order to increase the absorption efficiency, the detector can be arranged edge-on, in which case the absorption depth can be chosen to any length and the detector can still be fully depleted without going to very high voltages. For example, reference can be made to “Physical characterization of a scanning photon counting digital mammography system based on Si-strip detectors” by M. Åslund, B. Cederström, M. Lundqvist, and M. Danielsson in Med. Phys. 34, 2007, and “Two-View and Single-View Tomosynthesis versus Full-Field Digital Mammography: High-Resolution X-Ray Imaging Observer Study” by M G. Wallis, E. Moa, F. Zanca, K. Leifland and M. Danielsson in Radiology, 2012 Jan. 31.
U.S. Pat. No. 8,183,535 discloses an example of a photon-counting edge-on x-ray detector. In this patent, there are multiple semiconductor detector modules arranged together to form an overall detector area, where each semiconductor detector module comprises an x-ray sensor oriented edge-on to incoming x-rays and connected to integrated circuitry for registration of x-rays interacting in the x-ray sensor.
Photon-counting detectors based on direct conversion in simple semiconductor detectors, for example based on silicon or germanium, are well-known. For example, the use of a silicon diode for high-performance x-ray detectors is though less commonly used due to the more attractive features of e.g. CZT. A silicon diode is however cheaper to implement, but may require more signal processing.
Normally, the resolution is determined by the pitch of the diodes in one dimension and in the other dimension by the thickness of the wafer. Both the pitch for the diodes and the wafer thickness can be changed but to change the wafer thickness comes with a cost penalty and at some point, the wafers will be so thin that it will be hard or impossible to assemble the edge-on detectors without dead space between active areas.
Another possibility is to look at the time difference for the signal to reach the respective side of the wafers, the anode for the electrons and the cathode for the holes. This was studied and reported in “High resolution x-ray imaging using the signal time dependence on a double-sided silicon detector” by B. Cederström, M. Danielsson, M. Lundqvist and D. Nygren in Nucl. Instr. and Meth, 423 (1), pp. 135-145, 1999. The drawback with this arrangement is that it is impractical to combine the signals from the two sides of the wafer. The number of interconnects will increase dramatically.
The reference “Energy Calibration of a Silicon-Strip Detector for Photon-Counting Spectral CT by Direct Usage of the X-ray Tube Spectrum”, by Xuejin Liu, 2015 relates to a method by which a detector can be calibrated for variations in the silicon diode's physical properties by trimming the energy levels from a known source, and measure reference levels for comparison.
However, there is still room for improvements relating to the issue of increasing the resolution of photon-counting x-ray detectors.